The present disclosure relates generally to systems and methods for magnetic resonance imaging (MRI) and, in particular, to systems and methods for performing a spin-echo MRI process with rapid scan times, while still maintaining suitably a high signal-to-noise ratio (SNR) and suitably high-resolution images.
Any nucleus that possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant γ of the nucleus). Nuclei which exhibit this phenomena are referred to herein as “spins”.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment MZ is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, MZ, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides on signals which are emitted by the excited spins after the pulsed excitation signal B1 is terminated. Depending upon of biologically variable parameters such as proton density, longitudinal relaxation time (“T1”) describing the recovery of MZ along the polarizing field, and transverse relaxation time (“T2”) describing the decay of Mt in the x-y plane, this nuclear magnetic resonance (“NMR”) phenomena is exploited to obtain image contrast using different measurement sequences and by changing imaging parameters.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles that vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (Gx, Gy, and Gz) which have the same direction as the polarizing field B0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified. The acquisition of the NMR signals samples is referred to as sampling k-space, and a scan is completed when enough NMR cycles are performed to fully sample k-space.
One such process is referred to as the Fourier transform (FT) imaging technique, which is also referred to as “spin-warp” imaging. The spin-warp technique employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of NMR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional (“2D”) implementation, for example, spatial information is encoded in one direction by applying a phase encoding gradient (Gy) along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient (Gx) in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2D pulse sequence, the magnitude of the phase encoding gradient pulse Gy is incremented in the sequence of views that are acquired during the scan to produce a set of NMR data from which an entire image can be reconstructed.
In a three-dimensional (“3D”) implementation of the spin-warp method phase encoding of the spin-echo signals is performed along two orthogonal axes. In particular, a thick slab of spins is excited by applying a slab-selection gradient (Gz) in the presence of a selective RF excitation pulse and then a first phase encoding gradient (Gz) along the same axis and a second phase encoding gradient (Gy) are applied before the NMR signal acquisition in the presence of a readout gradient (Gx). For each value of the Gz phase encoding gradient, the Gy phase encoding is stepped through all its values to sample a three-dimensional region of k-space. By selectively exciting a slab, NMR signals are acquired from a controlled three-dimensional volume.
Commonly-used pulse sequences for generating T1, T2 and proton density-weighted imaging include “fast” spin-echo techniques, wherein a number of spin-echo signals forming a spin-echo train are generated due to multiple refocusing pulses following each radio-frequency (RF) excitation. For example, single-slab, 2D T2-weighted turbo spin echo (TSE) imaging techniques have been utilized to obtain high signal-to-noise ratio (SNR) and high resolution images, although such approaches do not fully utilize the scan time. Alternatively, multi-slab, 3D TSE imaging provides higher scan efficiency compared to 2D TSE, but is affected by ringing and venetian blind artifacts in the slice direction.
TSE-type sequences result in increased specific absorption rate (SAR) in tissue, which limits the utility of the technique due to FDA guidelines on power deposition, as well as magnetization transfer (MT) saturation, particularly at high polarizing fields, which therefore reduces imaging efficiency. Hence, various approaches have been proposed to help minimize the number of RF pulses needed and, thereby, the SAR, while maximizing efficiency and reducing artifacts. For example, some 3D TSE techniques have implemented spiral k-space trajectories, which provide advantages over conventional Cartesian approaches, including higher SNR efficiency, decreased scan time, and reduced number of RF pulses. In particular, as illustrated in FIGS. 1A and 1B, some readout strategies configure readout gradients to achieve uniform, or non-uniform, “spiral-in” or “spiral-out” trajectories for sampling k-space. Specifically, spiral methods typically sample k-space with an Archimedean or similar trajectory that begin at the k-space center and spiral to the edge (spiral-out), as illustrated in FIG. 1B, or its reverse, ending at the origin (spiral-in), as illustrated in FIG. 1A.
However, spiral imaging methods present additional complications and are often difficult to implement successfully, as images are typically subject to blurring and distortion caused by sensitivity to off-resonance and eddy current artifacts. For instance, as shown in FIG. 2, in many approaches, applied readout gradients 200 either do not align spin echoes 202 with the center of k-space, or only half of each echo is acquired if the spiral starts from the spin echo point (not shown).
Therefore, given the above shortcomings, there is a need for magnetic resonance imaging systems and methods that yield suitably high-SNR and suitably high-resolution within very rapid scan times.